Systems and methods for ablating body tissue

ABSTRACT

A transducer subassembly with combined imaging and therapeutic capabilities is disclosed. The subassembly includes heat sinks that are configured to maintain the transducer at a low operating temperature so that the transducer operates at high efficiency and also can handle a wider range of frequencies. The subassembly is also configured to allow cooling fluid to flow past the transducer element. One heat sink in the subassembly also acts as an acoustic matching layer and another heat sink acts as a backing Alternatively, the second heat sink which acts as a backing is optional. The transducer is configured to transmit at one power level for imaging, and at a second power level for ablating. The transducer may comprise sub-elements transmitting at different power levels. The subassembly may be operated at one power level for imaging and a second power level for ablating.

CROSS-REFERENCE

The present application is a continuation of U.S. application Ser. No.12/620,287 (Attorney Docket No. 31760-711.201) filed on Nov. 17, 2009,now U.S. Patent No. which is a non-provisional of, and claims thebenefit of priority of U.S. Provisional Patent Application No.61/115,403 (Attorney Docket No. 31760-711.101) filed Nov. 17, 2008, theentire contents of which are incorporated herein by reference.

The present application is related to U.S. Pat. No. 7,950,397 (AttorneyDocket No. 31760-703.201) and U.S. Pat. No. 7,942,871 (Attorney DocketNo. 31760-703.202) and copending U.S. patent application Ser. No.12/480,929 (Attorney Docket No. 31760-704.201); Ser. No. 12/480,256(Attorney Docket No. 31760-705.201); Ser. No. 12/483,174 (AttorneyDocket No. 31760-706.201); Ser. No. 12/482,640 (Attorney Docket No.31760-707.201); Ser. No. 12/505,326 (Attorney Docket No. 31760-708.201);Ser. No. 12/505,335 (Attorney Docket No. 31760-709.201); Ser. No.12/609,759 (Attorney Docket No. 31760-713.201); Ser. No. 12/609,274(Attorney Docket No. 31760-716.201); and Ser. No. 12/609,705 (AttorneyDocket No. 31760-718.201). The present application is also related tocopending U.S. Provisional Patent Application Nos. 61/148,809 (AttorneyDocket No. 31760-712.101); and 61/254,997 (Attorney Docket No.31760-720.101). The entire contents of each of the above referencedapplications is incorporated herein by reference.

BACKGROUND OF THE INVENTION 1. Field of the Invention

The present application generally relates to systems and methods forcreating ablation zones in human tissue. More specifically, the presentapplication relates to transducer configurations used to create tissuelesions, and even more specifically to ultrasound transducers used totreat fibrillation of the heart. While the present applicationemphasizes treatment of atrial fibrillation, one of skill in the artwill appreciate that this is not intended to be limiting, and that thesystems and methods disclosed herein may also be used to treat otherarrhythmias as well as to treating other conditions by creating lesionsin tissue.

The condition of atrial fibrillation (AF) is characterized by theabnormal (usually very rapid) beating of the left atrium of the heartwhich is out of synch with the normal synchronous movement (‘normalsinus rhythm’) of the heart muscle. In normal sinus rhythm, theelectrical impulses originate in the sino-atrial node (‘SA node’) whichresides in the right atrium. The abnormal beating of the atrial heartmuscle is known as ‘fibrillation’ and is caused by electrical impulsesoriginating instead at points other than the SA node, for example, inthe pulmonary veins (PV).

There are pharmacological treatments for this condition with varyingdegree of success. In addition, there are surgical interventions aimedat removing the aberrant electrical pathways from PV to the left atrium(‘LA’) such as the ‘Cox-Maze III Procedure’. This procedure has beenshown to be 99% effective but requires special surgical skills and istime consuming. Thus, there has been considerable effort to copy theCox-Maze procedure using a less invasive percutaneous catheter-basedapproach. Less invasive treatments have been developed which involve useof some form of energy to ablate (or kill) the tissue surrounding theaberrant focal point where the abnormal signals originate in PV. Themost common methodology is the use of radio-frequency (‘RF’) electricalenergy to heat the muscle tissue and thereby ablate it. The aberrantelectrical impulses are then prevented from traveling from PV to theatrium (achieving the ‘conduction block’) and thus avoiding thefibrillation of the atrial muscle. Other energy sources, such asmicrowave, laser, and ultrasound have been utilized to achieve theconduction block. In addition, techniques such as cryoablation,administration of ethanol, and the like have also been used. Some ofthese methods and devices are described below.

There has been considerable effort in developing catheter based systemsfor the treatment of AF using radiofrequency (RF) energy. One suchmethod includes a catheter having distal and proximal electrodes at thecatheter tip. The catheter can be bent in a coil shape, and positionedinside a pulmonary vein. The tissue of the inner wall of the PV isablated in an attempt to kill the source of the aberrant heart activity.

Another source used in ablation is microwave energy. One suchintraoperative device consists of a probe with a malleable antenna whichhas the ability to ablate the atrial tissue.

Still another catheter based method utilizes the cryogenic techniquewhere the tissue of the atrium is frozen below a temperature of −60degrees C. This results in killing of the tissue in the vicinity of thePV thereby eliminating the pathway for the aberrant signals causing theAF. Cryo-based techniques have also been a part of the partial Mazeprocedures described above. More recently, Dr. Cox and his group haveused cryoprobes (cryo-Maze) to duplicate the essentials of the Cox-MazeIII procedure.

More recent approaches for the treatment of AF involve the use ofultrasound energy. The target tissue of the region surrounding thepulmonary vein is heated with ultrasound energy emitted by one or moreultrasound transducers. One such approach includes a catheter distal tipportion equipped with a balloon and containing an ultrasound element.The balloon serves as an anchoring means to secure the tip of thecatheter in the pulmonary vein. The balloon portion of the catheter ispositioned in the selected pulmonary vein and the balloon is inflatedwith a fluid which is transparent to ultrasound energy. The transduceremits the ultrasound energy which travels to the target tissue in ornear the pulmonary vein and ablates it. The intended therapy is todestroy the electrical conduction path around a pulmonary vein andthereby restore the normal sinus rhythm. The therapy involves thecreation of a multiplicity of lesions around individual pulmonary veinsas required.

Yet another catheter device using ultrasound energy includes a catheterhaving a tip with an array of ultrasound elements in a grid pattern forthe purpose of creating a three dimensional image of the target tissue.An ablating ultrasound transducer is provided which is in the shape of aring which encircles the imaging grid. The ablating transducer emits aring of ultrasound energy at 10 MHz frequency.

While such ablation therapies alone are promising, it is preferred thatdevices and systems combine these ablation therapies with imagingcapabilities in a single unit. It would be particularly useful toprovide sensing or imaging (often used interchangeably) of the treatmentregion to properly position the ablation device relative to thetreatment region, as well as to evaluate progression of the treatment.Such imaging assists the system or the operator to ensure that only thetargeted tissue region is ablated. Furthermore, in a moving target suchas heart tissue, the original target identified by imaging, can move andthus non-target tissue may be inadvertently ablated. Hence,contemporaneous (or almost contemporaneous) imaging and ablationminimizes the risk of ablating non-target tissue. Thus, one unmet needusing ultrasound techniques for tissue ablation is to provide a devicecapable of both imaging as well as ablation.

Attaining this goal involves redesigning the key components of aconventional ultrasound ablation system to also provide an imagingfunction. Typically, ultrasound ablation is accomplished using atransducer assembly. The transducer assembly comprises a transducerelement, commonly one or more piezoelectrically active elements such aslead zirconate titanate (PZT) crystals. The PZT crystals often includean acoustical (impedance) matching layer on the ablating face tofacilitate efficient power transmission and to improve the imagingperformance. Further, the crystals may be bonded to a backing on thenon-ablative face to reflect or absorb any ultrasound beams in theappropriate direction. The conventional acoustic transducers which aretypically employed for the therapeutic purposes are acoustically large,often single-crystal devices having a narrower bandwidth in thefrequency domain than is required for good imaging performance. Althoughthey are designed to efficiently transmit acoustic energy to the targettissue, crystal devices with narrow bandwidth have previously beenviewed as unsuited for imaging. This has been due to the perceivedinability of conventional ablation transducers to handle the bandwidthof the ultrasound frequencies that would be optimized for both imagingand ablation. While ablation can be achieved using a narrower range offrequencies, imaging is usually performed using a wide range offrequencies. Thus, it is desirable that the PZT be able to accommodate awider bandwidth than used for ablation in order to accommodate theimaging bandwidth.

Wider transducer bandwidths are often achieved through the use ofmatching layers. Matching layers typically use materials with acousticimpedance between the acoustic impedances of the PZT and the tissue, andwith a thickness approaching ¼ wavelength of the ultrasound frequencyutilized. While matching layers are often used to improve thetransmission of ultrasound from the PZT into the tissue, they also canbe used to dampen the mechanical response of the PZT and broaden itsbandwidth. This dampening can result in some reduction of transducerefficiency. Furthermore wide bandwidth transducers may be unable operateat high power levels because they cannot be cooled effectively, partlydue to the thermally insulating properties of the matching layer. Aconventional PZT transducer with a higher bandwidth may often be only30%-50% efficient in converting the electrical energy to acousticenergy, and much of the energy is converted to heat and lost in thetransducer assembly. In addition to the lack of efficiency in convertingto ultrasound energy, the heat further reduces the PZT efficiency andmay cause the PZT crystal to depole and stop functioning as atransducer.

Thus, an additional challenge is to cool the transducer to maintain alower operating temperature than is presently provided for incommercially available systems. A cooled transducer can be drivenharder, i.e., it can tolerate higher electric powers and produce higheracoustic powers. This higher acoustic output is useful in increasing thelesion size and/or reducing the amount of time required to create alesion. Both of these attributes are important in the clinicalapplication of treating AF.

One method of cooling the transducer is to take advantage of the powerdensity and heat dissipation that are dependent on the size of thetransducer. As the diameter (and corresponding surface area) of thetransducer increases, the power density drops, and the heat dissipationper unit surface area also drops. If large enough, conventional coolingmethods may suffice to keep the transducer cool. However, in a cathetersuitable for ablation using an interventional approach, the transducermust necessarily be small and yet also be able to generate the powerdensity levels required to ablate tissue. In such a transducer, size isnot a suitable method of regulating the transducer's temperature. Thus,due to the small transducer size and consequent high power densities andlow heat dissipation, alternative approaches are warranted for coolingthe transducer.

One potential solution is the use of fluids to cool the transducer.Commonly, bodily fluids, such as blood flowing around the transducer,are used as a cooling fluid. However, blood tends to denature andcollect around the transducer when heated. In addition to the attendantproblems of possibly creating a clot in the atrium, the denatured bloodmay also adhere to the face of the transducer and create a layer ofinsulation, thereby further decreasing the performance of thetransducer. In contrast, introduced (non-bodily) fluids such as salineor water do not have the same attendant problems as blood and are usefulin maintaining lower transducer operating temperatures. However, inorder to be effective, these introduced fluids have to be effectivelytransported to the entire transducer to cool all the faces of thetransducer. If fluid transport is inadequate, the uncooled regions maydevelop “hot spots” that can impede the efficiency of the transducer.

While some devices, such as single crystal ultrasound therapy systemshave been reported for both imaging and therapeutic purposes, nonedisclose a method for cooling the entire transducer. Other multi-crystaltransducer assemblies are also available that circumvent the concerns ofthe single-crystal model. Some of these systems provide a method forcooling the back of the transducer crystals. However, none of thesesystems or methods include cooling of the entire transducer crystal. Asmentioned above, it is important to cool all the faces of the transducer(front and the back). Cooling only part of the transducer may lead to“hot spots” on some areas of the transducer, thereby decreasing theefficiency of the transducer in a situation where both ablation andimaging are necessary.

To realize combined imaging and ablation capabilities, some systems haveseparate imaging and ablation units. For example one commerciallyavailable system includes a treatment and imaging system. This systemcomprises a probe with an ultrasound transducer adapted to obtainimaging information from a patient treatment region, and also a separatearm member to deliver ultrasonic energy to the treatment region.Naturally, these are bulky and not well suited for use in catheter basedsystems. A variant of the combined imaging and ablation units is usingseparate transducer elements for imaging and ablation. This approachsuffers from many shortcomings including functionally, the ablatedtissue is not identical to the imaged tissue, and structurally thisconfiguration of discrete imaging and ablating elements occupies morespace in a housing, where space is limited in a transducer assembly,especially when the transducer is at the tip of the catheter as used inan interventional approach. Additionally, a multi-element device is moreexpensive and inconvenient to manufacture, along with the complicatedarrangements necessary for cooling the transducer elements. Further,multi-element devices are prone to misalignment, which may make themmore difficult to use. Also, multi-element devices typically requiremore complex and expensive systems for their control and use.

Thus, additional improvements are still desired in the field ofultrasound devices with combined imaging and ablating capabilities. Inparticular, it would be desirable to provide a device with asingle-crystal transducer assembly where all faces of the transducercrystal are cooled to protect and preserve the operating efficiency. Itwould also be desirable to provide such a system that is easy to use,easy to manufacture and that is lower in cost than current commercialsystems.

2. Description of Background Art

Patents related to the treatment of atrial fibrillation include, but arenot limited to the following: U.S. Pat. Nos. 7,393,325; 7,142,905;6,997,925; 6,996,908; 6,966,908; 6,964,660; 6,955,173; 6,954,977;6,953,460; 6,949,097; 6,929,639; 6,872,205; 6,814,733; 6,780,183;6,666,858; 6,652,515; 6,635,054; 6,605,084; 6,547,788; 6,514,249;6,502,576; 6,500,121; 6,416,511; 6,383,151; 6,305,378; 6,254,599;6,245,064; 6,164,283; 6,161,543; 6,117,101; 6,064,902; 6,052,576;6,024,740; 6,012,457; 5,629,906; 5,405,346; 5,314,466; 5,295,484;5,246,438; 4,757,820 and 4,641,649.

Patent Publications related to the treatment of atrial fibrillationinclude, but are not limited to International PCT Publication Nos. WO2005/117734; WO 1999/002096; and U.S. Patent Publication Nos.2005/0267453; 2003/0050631; 2003/0050630; and 2002/0087151.

Scientific publications related to the treatment of atrial fibrillationinclude, but are not limited to: Haissaguerre, M. et al., SpontaneousInitiation of Atrial Fibrillation by Ectopic Beats Originating in thePulmonary Veins, New England J Med., Vol. 339:659-666; J. L. Cox et al.,The Development of the Maze Procedure for the Treatment of AtrialFibrillation, Seminars in Thoracic & Cardiovascular Surgery, 2000; 12:2-14; J. L. Cox et al., Electrophysiologic Basis, Surgical Development,and Clinical Results of the Maze Procedure for Atrial Flutter and AtrialFibrillation, Advances in Cardiac Surgery, 1995; 6: 1-67; J. L. Cox etal., Modification of the Maze Procedure for Atrial Flutter and AtrialFibrillation. II, Surgical Technique of the Maze III Procedure, Journalof Thoracic & Cardiovascular Surgery, 1995; 110:485-95; J. L. Cox, N.Ad, T. Palazzo, et al. Current Status of the Maze Procedure for theTreatment of Atrial Fibrillation, Seminars in Thoracic & CardiovascularSurgery, 2000; 12: 15-19; M. Levinson, Endocardial Microwave Ablation: ANew Surgical Approach for Atrial Fibrillation; The Heart Surgery Forum,2006; Maessen et al., Beating Heart Surgical Treatment of AtrialFibrillation with Microwave Ablation, Ann Thorac Surg 74: 1160-8, 2002;A. M. Gillinov, E. H. Blackstone and P. M. McCarthy, AtrialFibrillation: Current Surgical Options and their Assessment, Annals ofThoracic Surgery 2002; 74:2210-7; Sueda T., Nagata H., Orihashi K., etal., Efficacy of a Simple Left Atrial Procedure for Chronic AtrialFibrillation in Mitral Valve Operations, Ann Thorac Surg 1997;63:1070-1075; Sueda T., Nagata H., Shikata H., et al.; Simple LeftAtrial Procedure for Chronic Atrial Fibrillation Associated with MitralValve Disease, Ann Thorac Surg 1996; 62:1796-1800; Nathan H., EliakimM., The Junction Between the Left Atrium and the Pulmonary Veins, AnAnatomic Study of Human Hearts, Circulation 1966;34:412-422; Cox J. L.,Schuessler R. B., Boineau J. P., The Development of the Maze Procedurefor the Treatment of Atrial Fibrillation, Semin Thorac Cardiovasc Surg2000;12:2-14; and Gentry et al., Integrated Catheter for 3-DIntracardiac Echocardiography and Ultrasound Ablation, IEEE Transactionson Ultrasonics, Ferroelectrics, and Frequency Control, Vol. 51, No. 7,pp 799-807.

SUMMARY OF THE INVENTION

The present invention discloses a transducer assembly with combinedimaging and therapeutic capabilities that may be used to create lesionsin tissue. In preferred embodiments, the transducer assembly is used toablate tissue to create a conduction block in the target tissue whichblocks aberrant electrical pathways. Thus, the transducer assembly maybe used as a treatment for fibrillation or other arrhythmias, as well asother conditions requiring creation of a lesion in tissue.

In a first aspect of the present invention, a transducer systemcomprises a transducer element comprising a proximal surface and adistal surface, and a first heat sink attached to the distal surface ofthe transducer element. The system also has a second heat sink attachedto the proximal surface of the transducer element, and a base coupled tothe first and second heat sinks The base is configured to allow fluidflow past the transducer element for cooling of the and distal surfacesof the transducer element, and the heat sinks

The system may further comprise a tubular jacket configured to house thebase, the transducer element, and the first and second heat sinks Thetubular jacket may comprise at least one fluid exit port configured toallow fluid to exit the tubular jacket. The first heat sink may comprisea first bonding portion and a first substantially bent portion. Thefirst bonding portion may be bonded to the distal surface of thetransducer, and the first substantially bent portion may protrudeproximally from the transducer element, thereby conducting heat awayfrom the distal surface of the transducer element. The first bondingportion may comprise a material whose composition and dimension providesan acoustically matching layer on the distal surface of the transducerelement. The first bonding portion may comprise a material chosen fromthe group consisting of aluminum, graphite, metal-filled graphite,ceramic, an amalgam of graphite and copper or tungsten, and anepoxy-filled metal. The bonding portion may be in electrical and/orthermal communication with the distal surface of the transducer element.Electrical communication between the bonding portion and the distalsurface may be established by direct contact between the bonding portionand the distal surface. The direct contact may be controlled by surfaceroughness of the bonding portion and the distal surface.

The second heat sink may comprise a second bonding portion and a secondsubstantially bent portion. The second bonding portion may be bonded tothe proximal surface of the transducer, and the second substantiallybent portion may protrude proximally from the transducer element,thereby conducting heat away from the proximal surface of the transducerelement. The second bonding portion may comprise a material whosecomposition is acoustically mismatched to an acoustic impedance of thetransducer element, thereby providing a reflective backing layer on theproximal surface of the transducer element. The second bonding portionmay comprise a metal such as copper. An air pocket may be disposedbetween the proximal surface of the transducer and the second heat sink.

The transducer element may comprise a substantially flat circular disc,and the transducer element may operate at a first power level in a firstfrequency range and a second power level in a second frequency range.The first frequency range may be used for ultrasonically imaging tissueand the second frequency range may be used for creating tissue lesions.The first frequency range may be 5 MHz to 30 MHz and the secondfrequency range may be 10 to 18 MHz.

The first and second bonding portions may comprise a matrix containingperforations such that the first bonding portion is acoustically matchedand the second bonding portion is acoustically mismatched to theacoustic impedance of the transducer element. The system may furthercomprise an elongate flexible shaft having a proximal end and a distalend, and the transducer may be disposed adjacent the distal end of theshaft. The system may also comprise a cooling fluid in fluidcommunication with the transducer. The system may comprise a temperaturesensor adjacent the transducer for monitoring temperature of thetransducer or cooling fluid flowing therepast. Adjustments to thecooling fluid flow rate or the transducer power levels may be made basedon the monitored temperature.

In another aspect of the present invention, a method of ablating tissuecomprises introducing an ablation device into a patient. The devicecomprises an ultrasound transducer element configured to operate at afirst power level and at a second power level. The first power level isused for ultrasonically imaging tissue and identifying a target tissue,and the second power level is used for ablating the target tissue.Operating the transducer element at the first power level allows imagingof a portion of the tissue and identification of the target tissue.Operating at the second power level ablates the target tissue. Theultrasound transducer surfaces are cooled during operation.

The transducer element may comprise a proximal surface and a distalsurface, and the device may further comprise first and second heat sinksbonded to the distal and proximal surfaces of the transducer element,respectively. The cooling step may comprise introducing fluid to thetransducer element and to the first and second heat sinks duringoperation of the transducer element, thereby further cooling thetransducer element. The transducer element may comprise first and secondportions. The first portion may be configured to operate at the firstpower level and the second portion may be configured to operate at thesecond power level. The first portion may be operated at the first powerlevel concurrently with operation of the second portion at the secondpower level. The introducing step may comprise passing the ablationdevice transseptally across a septal wall of the patient's heart. Theintroducing step may also comprise positioning the ablation device intoa left atrium of the patient's heart. There may not be direct contactbetween the transducer and the target tissue.

These and other embodiments are described in further detail in thefollowing description related to the appended drawing figures.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A illustrates an exemplary system for treating tissue using atransducer assembly.

FIGS. 1B-1C illustrate exemplary embodiments of a transducer assembly.

FIGS. 2A-2D illustrate alternative embodiments of the transducerelement.

FIG. 3 illustrates the transducer element with a first heat sink.

FIG. 4 illustrates the transducer element with a second heat sink.

FIG. 5 illustrates the transducer assembly in a tubular jacket.

FIG. 6 illustrates an ablation pattern in tissue.

FIGS. 7A-7D illustrate the progression of ablation in tissue.

FIG. 8 illustrates an alternative lesion shape.

DETAILED DESCRIPTION OF THE INVENTION

Although the detailed description contains many specifics, these shouldnot be construed as limiting the scope of the invention but merely asillustrating different examples and aspects of the invention. It shouldbe appreciated that the scope of the invention includes otherembodiments not discussed in detail above. Various other modifications,changes and variations which will be apparent to those skilled in theart may be made in the arrangement, operation and details of the methodand apparatus of the present invention disclosed herein withoutdeparting from the spirit and scope of the invention as described here.

The present invention relates to creating ablation zones in humantissue, and more specifically to transducer assemblies (orsubassemblies) that are used for creating tissue lesions. FIG. 1A is adiagrammatic illustration of an exemplary embodiment of a system forcreating ablation zones in human tissue, as described in the abovereferenced related parent applications. A catheter device C is housedwithin a sheath S. A proximal portion of the catheter C is coupled to aconsole P. A distal portion of the catheter C, comprising an ultrasonictransducer subassembly T, is introduced into the heart, preferablytransseptally, into the left atrium (LA), adjacent the pulmonary veinsPV of a patient. The transducer subassembly T is energized to provideultrasonic energy for ablating tissue. The console P controls energydelivery to the transducer subassembly T, as well as movements of thedistal portion of the catheter C to trace ablation paths. Additionaldetails on the ablation system are disclosed in U.S. Provisional PatentApplication No. 61/254,997 previously incorporated herein by reference.

For brevity, the transducer subassemblies are described herein withrespect to one embodiment of a catheter for sensing and ablating tissue.However, the transducer assemblies of this invention may be utilizedwith any suitable device in both medical and non-medical fields.

The transducer subassemblies comprise transducer elements and areconfigured such that the same transducer element may be used to bothimage (for example, in A-mode) and ablate. The transducer elements maybe in the shape of a disc, or other shapes may be used for thetransducer elements. The transducer subassemblies are also configuredfor effective cooling of the transducer elements, in order to increasethe efficiency of transduction. This is accomplished by affixing (e.g.by bonding, welding, snap fitting, etc.) a distal and a proximal heatsink to the transducer element, thereby conducting heat away from thetransducer element. In order to further increase efficiency, the distalheat sink comprises an acoustically matching layer and the proximal heatsink comprises an acoustically mismatched backing layer. Additionally,each of the heat sinks is configured to allow for a cooling substance(e.g., a fluid such as saline, water) to be directed to and dissipatethe heat from the proximal and distal surfaces (hereinafter alsoreferred to as “faces”) of the transducer element.

As shown in FIG. 1B, a transducer subassembly 3000 is placed at or nearthe distal portion of a catheter 2000 and contained within a tubularjacket 3400. The catheter 2000 may be any suitable catheter andcomprises at least one lumen 2100. The components of transducersubassembly 3000 are shown in an assembled view in FIG. 1B, and in anexploded view in FIG. 1C. The transducer subassembly 3000 comprises atransducer element 3100 having a distal face 3102 and a proximal face3104. The transducer subassembly 3000 further comprises heat sinks thatserve to cool the transducer element 3000 by conducting heat away fromit. Specifically, the transducer subassembly 3000 comprises a distalheat sink 3300 bonded to the distal face 3102 of the transducer element3100, and a proximal heat sink 3200 bonded to the proximal face 3104 ofthe transducer element 3100.

The heat sinks are further configured to increase the operatingefficiency of the transducer element 3000 through acoustic matching andacoustic reflection. Specifically, and as described in further detailbelow, the distal heat sink 3300 comprises an acoustically matchinglayer portion, i.e., a portion whose composition and thickness providesa ¼ wavelength matching layer between the transducer element 3100 andany fluid in front of the transducer subassembly 3000. The proximal heatsink 3200 comprises an acoustically mismatched layer portion, i.e., aportion whose composition is acoustically mismatched to the acousticimpedance of the transducer element 3100, thereby reflecting ultrasoundwaves emanating from the transducer element 3100 back towards thetransducer element 3100. These portions are more fully described below.

The transducer subassembly 3000 also comprises a base 3500 anchoring theheat sinks 3200 and 3300, with the transducer element 3100 bondedbetween the heat sinks The transducer subassembly 3000 is powered usingone or more electrical cables 3600 bonded to each of the heat sinks 3200and 3300. These electrical cables 3600 are exemplarily provided througha pair of twisted wires, as shown in FIGS. 1B and 1C. As will beappreciated, they could also be coaxial or separate untwisted wires. Theheat sinks 3200 and 3300 comprise electrical attachments (not shown) forelectrically coupling the heat sinks 3200 and 3300 to the electricalcables 3600, thereby providing electrical power to the transducerelement 3100. The transducer element 3100 comprises electrode platingson the distal and proximal faces in order to distribute the electricalenergy over the faces of the transducer element 3100.

As disclosed herein, the transducer element 3100 comprises a singletransducer element. However, those skilled in the art would appreciatethat this single element may be comprised of smaller sub-elements. Thetransducer is of a suitable size to fit into a catheter configured to beintroduced percutaneously into the atria of the heart. For example, inone embodiment, the transducer diameter is less than 0.2 inches, andpreferably less than 0.15 inches.

Further, the transducer element may comprise a variety of geometries, aswell as a variety of acoustically active and inactive portions. Suchtransducer element properties in turn influence the transducer's imagingand ablative properties, such as the shape of the created ablationlesions. These concepts of using transducer elements of various shapesand sizes (sub-elements) are further described below.

For example, in the embodiment shown in FIGS. 1B and 1C, the transducerelement 3100 is a flat, circular disc that transmits ultrasound energyfrom its proximal and distal faces. The transducer element 3100 mayalternatively have more complex geometry, such as either concave orconvex, to achieve an effect of a lens or to assist in apodization(i.e., in selectively decreasing the vibration of a portion or portionsof the surfaces of the transducer element 3100) and management of thepropagation of the ultrasound beam.

Other exemplary transducers are shown in FIGS. 2A through 2D. Forexample, as shown in FIGS. 2A and 2B, the transducers 3100 a and 3100 binclude at least one acoustically inactive portion 4200, with theremainder of the transducer surface comprising an acoustically activeportion. In these embodiments, the acoustically inactive portion 4200does not emit an energy beam when the transducer is energized, or mayalternatively emit an energy beam with a very low (substantially zero)energy. The acoustically inactive portion 4200 has several functions.For instance, the shape of a lesion produced by ablating tissue usingsuch a transducer may correspond with the shape of the acousticallyactive ablating portions. For example, in the circular embodiment shownin FIGS. 1B and 1C, the shape of the lesion will be tear-drop shaped.However, in the annular embodiment shown in FIG. 2A, the shape of thelesion will be approximately tooth-shaped or a blunted tear-shaped. Thisis because the acoustically inactive portion 4200 in FIG. 2A willpreclude prolonged ablation at the corresponding central portion of thetissue. Since prolonged ablation of tissue creates a deeper ablation,the presence of acoustically inactive portion 4200 precludes ablationfrom reaching further into the tissue at the central portion. The lesionthus is approximately tooth-shaped or blunted tear-shaped, asillustrated by the exemplary lesion shape L of FIG. 2A, rather thantear-shaped.

In addition to influencing the shape of the created ablation lesion,acoustically inactive portion 4200, in any of the embodiments shown,further functions to aid in the temperature regulation of the transducerelements 3100 a and 3100 b, i.e., in preventing the transducer elementsfrom becoming too hot.

Acoustically inactive portions may be created in a variety of ways. Inone embodiment, an acoustically inactive portion 4200 is a hole or gapdefined by the boundary of the acoustically active region of thetransducer element. In such an embodiment, an optional coolant sourcemay be coupled to (or in the case of a coolant fluid, it may flowthrough) the hole or gap defined by the transducer element to furthercool and regulate the temperature of the transducer element.

In another embodiment, the acoustically inactive portion 4200 maycomprise a material composition with different properties from that ofthe active region of the transducer element. For example, theacoustically inactive material may be made of a metal, such as copper,which further functions to draw or conduct heat away from the transducerelement. Alternatively, the acoustically inactive portion 4200 may bemade from the same material as the transducer element, but with theelectrode plating removed or disconnected from the electricalattachments. The acoustically inactive portion 4200 may be disposedalong the full thickness of the transducer element, or may alternativelybe a layer of material on or within the transducer element that has athickness less than the full thickness of the transducer element.

For example, as shown in FIG. 2A, the transducer element 3100 a is adoughnut-shaped transducer that comprises a hole (or acousticallyinactive portion) 4200 in the center portion of the otherwise circulardisc-shaped transducer element. The transducer element 3100 a of thisembodiment has a circular geometry, but may alternatively be elliptical,polygonal as shown in FIG. 2B, or any other suitable shape. Thetransducer element 3100 a includes a singular, circular acousticallyinactive portion 4200, but may alternatively include any suitable numberof acoustically inactive portions 4200 of any suitable geometry, asshown in FIG. 2B. Exemplary geometries of acoustically inactive portionsinclude circular, square, rectangular, elliptical, polygon, or any othershaped region. The total energy emitted from the transducer element isrelated to the acoustically active surface area of the transducerelement. Therefore, the size and location of acoustically inactiveportion(s) 4200 may sufficiently reduce the heat build-up in thetransducer element, while allowing the transducer element to provide asmuch output energy as possible or as desired.

As disclosed herein, the transducer elements may optionally beconfigured to operate at more than one frequency. This allows them to beused for multi-frequency ablating or for contemporaneous ablation anddiagnosis. For example, such a multi-frequency transducer element may beoperated intermittently at a first power level using a first frequencyrange that is used to image a portion of the tissue in order to identifya target tissue, and operated at a second power level using a secondfrequency range that is used to ablate the target tissue. In oneembodiment, the imaging frequency is in the range of about 5 MHz to 30MHz, and the ablation frequency is preferably in the range of 5 to 25MHz, more preferably in the range 8 to 20 MHz, and even more preferablyin the range 10 to 18 MHz. The transducers achieving theseconfigurations are shown to exemplarily be annular transducers or gridarrays.

As shown in FIGS. 2C and 2D, the transducer elements 3100 c and 3100 dare configured to be capable of transmitting at more than one frequency.Specifically, as shown in FIG. 2C, the transducer element 3100 cincludes a plurality of annular transducer portions 4400. The pluralityof annular transducer portions is a plurality of concentric rings, butmay alternatively have any suitable configuration with any suitablegeometry, such as elliptical or polygonal. Optionally, the transducerelement 3100 c includes one or more acoustically inactive portions 4200,such as the center portion of the transducer 3100 c. The plurality ofannular transducer portions 4400 includes at least a first annularportion and a second annular portion. The first annular portion may havematerial properties that differ from those of the second annularportion, such that the first annular portion emits a first energy beamthat is different from a second energy beam emitted by the secondannular portion. Furthermore, the first annular portion may be energizedwith a different frequency, voltage, duty cycle, power, and/or for adifferent length of time from the second annular portion. Alternativelythe first annular portion may be operated in a different mode from thesecond annular portion. For example, the first annular portion may beoperated in a therapy mode, such as ablation mode, which delivers apulse of ultrasound energy sufficient for heating the tissue. The secondannular portion may be operated in an imaging mode, such as A-mode,which delivers a pulse of ultrasound of short duration, which isgenerally not sufficient for heating of the tissue but functions todetect characteristics of the target tissue and/or environment in andaround the ultrasound delivery system. The first annular portion mayfurther include a separate electrical attachment from that of the secondannular portion.

In a another embodiment of a multi-frequency transducer element shown inFIG. 2D, the transducer element 3100 d includes a grid of transducerportions 4600. The grid of transducer portions 4600 has any suitablegeometry such as circular, rectangular, elliptical, polygonal, or anyother suitable geometry. The transducer element 3100 d in this variationmay further include one or more transducer portions that areacoustically inactive. The grid of transducer portions 4600 includes atleast a first transducer portion and a second transducer portion. Thefirst transducer portion and the second transducer portion are portionsof a single transducer with a single set of material properties. Thefirst transducer portion is energized with a different frequency,voltage, duty cycle, power, and/or for a different length of time fromthe second transducer portion. Furthermore, the first transducer portionmay be operated in a different mode from the second transducer portion.For example, similar to the description above, the first transducerportion may operate in a therapy mode, such as ablate mode, while thesecond transducer portion may operate in a imaging mode, such as A-mode.The first transducer portion may further include a separate electricalattachment from that of the second transducer portion. For example, thefirst transducer portion may be located towards the center of thetransducer element 3100 d and the second transducer portion may belocated towards the outer portion of the transducer element 3100 d.Further, the second transducer portion may be energized while the firsttransducer portion remains inactive. In other embodiments, the firsttransducer portion has material properties that differ from those of thesecond transducer portion, such that the first transducer portion emitsa first energy beam that is different from a second energy beam emittedfrom the second transducer portion. In such an embodiment, the firsttransducer portion may also be energized with a different frequency,voltage, duty cycle, power, and/or for a different length of time fromthe second transducer portion.

Turning now to the heat sinks 3200 and 3300, FIG. 3 shows the proximalheat sink 3200. In this embodiment, the proximal heat sink 3200comprises a bonding portion 3210 and a substantially bent portionforming legs 3220 that are generally orthogonal to the bonding portion3210. The proximal heat sink further comprises at least one electricalattachment 3230. Similarly, the distal heat sink comprises an electricalattachment 3330 (shown in FIG. 4). The electrical wires 3600 areconnected to the electrical attachments 3230 and 3330. Unlikeconventional electrical attachments to a transducer crystal, where theelectrical leads are connected to the opposing faces of the crystal, thedisclosed arrangement eliminates “hot spots” and results in a uniformelectrical power density across the surface of the crystal.Additionally, this results in an easier assembly or manufacturingprocess.

The bonding portion 3210 is bonded to the proximal face of thetransducer element 3100 with a suitable bonding material such as anepoxy to form a bond layer. Though shown as substantially flat in thisembodiment, one skilled in the art will appreciate that the bondingportion 3210 may be any suitable configuration such as a concave portionto still maintain the functionality described herein. The substantiallybent portion 3220 comprises legs, or elements that protrude proximallyfrom the transducer element 3100. Further, the bent portion 3220 isconfigured in a manner to allow for fluid to flow through the bentportion and also allows the fluid to surround and cool the proximal faceof the transducer element 3100. The fluid that could be accommodatedwithin the bent portion could be any suitable fluid that achieves anappropriate balance between having an effective heat sink and minimizingacoustic reverberations that degrade image performance. The proximalheat sink 3200 is formed from a suitable material such as copper of asuitable thickness. The thickness of the material for this heat sinkpreferably ranges between 0.0001 inches to 0.01 inches for a copper heatsink.

Proximal heat sink 3200 serves to cool the proximal face of thetransducer by conducting and dissipating the heat away from thetransducer element 3100. Heat from the transducer element 3100 isabsorbed by the bonding portion 3210, and conducted to the bent portion3220 where it is dissipated into the circulating fluid. This dissipationprovides some cooling to the proximal face of the transducer element3100. Additionally, the bent portion 3220 is configured in a manner toallow for fluid to surround and cool the proximal face of the transducerelement 3100. For example, as shown in FIG. 3, the bent portion 3220provides for one or more pockets behind the transducer element 3100where a fluid may be introduced to flow and cool both the transducerelement 3100 as well as the proximal heat sink 3200 that has dissipatedheat from the proximal face of the transducer element 3100.

As described above, in addition to dissipating the heat, the proximalheat sink 3200 also serves as a heat spreader to reduce hot spots in thetransducer element 3100, and thereby preserve it over its entire face.Without this heat spreading, the center of the transducer element 3100would be substantially hotter than the rest of the transducer element3100.

The bonding portion 3210 can be configured to maximize the amount ofreflected energy transmitted from the transducer element 3100. Sincemany metals suitable for heat sink applications have acoustic impedancesthat are not too dissimilar from PZT, the boundary between PZT and theheat sink itself does not provide a very effective reflective interface.However, another material immediately proximal to the heat shield couldbe selected so that it provides an efficient acoustic reflector. Forexample, air provides an excellent acoustic mismatch, as does water, andtherefore acts as good reflectors. Water is preferred since it also actsas a thermal conductor, even though it is not quite as effective areflector as air. Air could be used, provided that it does not interferewith the flow of cooling fluid around the transducer assembly. Toaccomplish this, the bonding portion of 3210 could be constructed fromtwo metal layers capturing a third thin layer of air in between.Alternatively, a backing material may be located proximal to theproximal heat sink 3200 to provide an acoustically absorptive medium tominimize reverberations to further optimize imaging performance. Suchbacking materials may optionally be made of combinations of epoxy, metalparticles, tungsten and the like.

Additionally or alternatively, the transducer element 3100 or thetransducer subassembly 3000 may be placed on a tripod-style structure(not shown) such that the proximal surface of the transducer element3100 faces into the tripod. In this configuration, a pocket forms in thespace between the transducer element 3100 and the tripod base. Thispocket serves as an alternative backing with the same two-fold purpose.First, it is acoustically mismatched and thereby reflective of theultrasound waves emanating from the transducer element 3100. Second, asfluid (for example saline or water) is introduced into the transducerassembly 3000, the pocket also allows for the fluid to come into contactwith the transducer element 3100 and thereby provide for additionalcooling.

Alternatively, another suitable acoustically mismatched material withreasonable thermal conduction could be used in place of fluid. Suchmaterials include metal with trapped air, for example steel wool orporous metal with entrapped air. For example, the rear of the PZT maycomprise a thin heat spreader comprising the entire rear face with apocket of porous metal attached behind. As another example, the centerof the PZT could be further cooled by providing a thermally conductingcenter post as part of the heat sink, allowing an annular ring of air tobe trapped behind the bonding portion 3210.

As mentioned above, additional cooling can be provided by a distal heatsink 3300 (which also serves as a heat spreader) for distributing theheat and cooling the distal face of the transducer element 3100. Asshown in FIG. 4, the distal heat sink 3300 also comprises a bondingportion 3310 and a substantially bent portion 3320 that is orthogonal tothe flat portion 3310. The distal heat sink further comprises at leastone electrical attachment 3330. The distal heat sink 3300 is configuredsuch that the bonding portion 3310 is bonded to the distal face of thetransducer element 3100. The substantially bent portion 3320 compriseselements or legs that protrude proximally from the transducer element3100. Thus the bent portion 3320 of the distal heat sink 3300 isadjacent to the bent portion 3220 of the proximal heat sink 3200. Asmentioned above, the bonding portion 3310 is further configured to serveas an acoustically matching layer for the transducer element 3100. Toprovide an acoustically matching composition that is also thermallyconductive, the bonding portion 3310 is made of a suitable material suchas aluminum; other such suitable materials include graphite,metal-filled graphite or ceramic, or an amalgam of graphite and copperor tungsten, in suitable thickness that range from 0.026 inches to0.00026 inches so that it is ¼ wavelength at the desired frequency. Thebonding portion 3310 is bonded to the distal face of the transducerelement 3100 with a suitable bonding material such as an epoxy to form abond layer.

Additionally and optionally, the bonding portion 3310 comprisesperforations or holes 3315 that may be filled with epoxy applied in alayer of a suitable thinness to enhance the acoustic impedance matching.Perforations in the distal matching layer can be accomplished in manyways. The perforated structure is made of a combination of metal matrixcontaining open spaces, later to be filled with an epoxy material. Forexample, the metal matrix can be a wire grid. Alternatively, theperforated structure may be a matrix of epoxy film, and the holes may befilled with a metal such as aluminum. Additionally, the ratio of epoxyto the metal mixture is configured to enhance acoustic impedancematching. The acoustic impedance is determined by the acoustic impedanceof the two composite materials, and the ratio of the mixture. Forexample, using aluminum and EPO-TEK® 377 (Epoxy Technology, Inc.,Billerica, Mass.) the appropriate ratio is 35-60% volume fraction ofepoxy and a good acoustic impedance matching is achieved at a 40-50%volume fraction of epoxy and an ideal match about 41%. Additionally, theperforations or holes 3315 have a sufficiently small diameter ascompared to the wavelength of the ultrasonic beam, thereby allowing thebonding portion 3310 to appear homogeneous to the propagating wavesemanating from the transducer element 3100.

Similar to the construction of using bonding portion 3310 withperforations or holes to achieve acoustic impedance matching, thebonding portion 3210 at proximal surface of the transducer crystal alsomay benefit from using perforations or holes in the material used toachieve acoustic impedance mismatch. Such materials may include copper,tungsten and the like. Alternatively, an epoxy layer with metalparticles sprinkled in it and a distribution of holes or perforationsmay achieve the same purpose of providing acoustic impedance mismatch.

Both non-conductive and conductive epoxy (with metal particles such assilver) could be used to form either the proximal or distal bond layer.In one embodiment, the epoxy is exemplarily a non-conductive epoxy of alow viscosity (e.g., EPO-TEK® 377). The epoxy is applied in a layer ofsuitable thinness to minimize its impact on acoustic impedance matching,while maximizing thermal conduction to cool the transducer 3100.Additionally, the bond layers are also configured to electricallyconnect the heat sinks 3310 and 3210 to the transducer 3100. This issuccessfully accomplished without the use of conductive epoxy byconfiguring the transducer 3100 faces and the bonding portions 3310 and3210 to be rough. Thereafter, the distal and the proximal faces of thetransducer element 3100 are bonded to their relevant heat sinks withelectrically non-conductive epoxy. Each bond layer is of sufficientthinness to allow the surface roughness of the transducer 3100 toelectrically contact the surface roughness of the heat sinks 3310 and3210. This allows the rough surfaces of the transducer element 3100 tocome into direct electrical contact with their relevant heat sinks,thereby obviating the need for using electrically conductive epoxy(which may degrade with heat). Thus, electrical conduction occurs viathe contact points between the rough surfaces of the transducer element3100 and the heat sinks, rather than through the epoxy.

Additionally and optionally, parylene or any such suitable coating isdisposed on the bonding portion 3310 of the distal heat sink 3300 to actas an additional matching layer. One result of the coating may be tothus produce a second acoustic matching layer for increased efficiencyof transducer element 3100 conduction and to further optimize the widebandwidth performance. The thickness of this parylene coat is ¼ of thetarget ultrasound wavelength. Optionally, both heat sinks 3200 and 3300are coated with parylene or any such suitable coatings to provideelectrical isolation. Further, heat sinks are anodized to provideelectrical isolation while maximizing thermal conduction. The transducersubassembly 3000 is located within a tubular jacket 3400, as shown inFIG. 5. The tubular jacket 3400 is a hollow cylinder with a proximal anddistal end. The transducer subassembly 3000 is placed into the tubularjacket 3400 such that the distal end of the tubular jacket protrudes asuitable distance, for example between 1 mm to 5 mm beyond the distalend of the transducer subassembly 3000. The distal end of the tubularjacket 3400 comprises a distal opening 3410, and fluid exit ports 3420located near the distal opening. Cooling of the transducer element 3100may be accomplished by introducing a cooling fluid or gel, such assaline, water, or any physiologically compatible fluid or gel, into theproximal end of the tubular jacket 3400. The cooling fluid has a lowertemperature relative to the temperature of the transducer element 3100.The cooling fluid flows along the bent portions 3220 and 3320 of heatsinks 3200 and 3300 and over both bonding portions 3210 and 3310 andexits through the distal opening 3410, the fluid exit ports 3420, or anycombination thereof. Optionally, the exit ports 3420 may be in the formof a grating, a screen, holes, drip holes, a weeping structure, or anyof a number of suitable apertures.

Additionally, any or all of the metal components described in transducersubassembly 3000 are provided with a plating of a suitable biocompatiblematerial such as gold. Such plating is provided to the individualcomponents before the transducer assembly is assembled.

In an exemplary embodiment, the temperature of the cooling fluid or gelis sufficiently low that it cools the transducer element 3100 and,optionally, the target tissue. In this embodiment, the temperature ofthe fluid or gel is between approximately −5 and body temperature. In asecond embodiment, the temperature of the cooling fluid or gel is withina temperature range such that it cools the transducer element 3100, butdoes not cool the target tissue, and may actually warm the targettissue. The fluid or gel may alternatively be any suitable temperature,including room temperature, to sufficiently cool the transducer element3100.

The invention described above has the advantage of keeping the smallertransducer assembly cool. As previously mentioned, the transducerdiameter is small enough (less than 0.2 inches, and ideally less than0.15 inches) to fit into the tip of a catheter and yet generate powerdensity levels that are high enough to create tissue lesions (about 50watts/cm² to 2500 watts/cm²). This invention keeps the transducerassembly cool in order to create tissue lesions efficiently.

We now turn to describing the formation of lesions. The interaction ofthe ultrasound beam with the tissue is shown in FIG. 6. The tissue 276is presented to the ultrasound beam 272 within a collimated length L.The front surface 280 of the tissue 276 is at a distance d (282) awayfrom the distal tip 2110 of the catheter 2000. As the ultrasound beam272 travels through the tissue 276, its energy is absorbed and scatteredby the tissue 276, and most of the ultrasound energy is converted tothermal energy. This thermal energy heats the tissue to temperatureshigher than the surrounding tissue. The result is a heated zone 278which has a typical shape of an elongated tear drop. The diameter D1 ofthe zone 278 is smaller than the transducer aperture diameter D at thetissue surface 280, and further, the outer layer(s) of tissue 276 remainsubstantially undamaged. This is due to the thermal cooling provided bythe surrounding fluid which is flowing past the tissue surface 280. Moreor less of the outer layers of tissue 276 may be spared or may remainsubstantially undamaged, depending on the amount that the tissue surface280 is cooled and/or depending on the characteristics of the ultrasounddelivery system (including the transducer element 3100 the ultrasoundbeam 272, the ultrasound energy and the frequency). The energy depositedin the ablation zone 278 interacts with the tissue such that theendocardial surface remains pristine and/or not charred. As theultrasound beam 272 travels deeper into the tissue 276, thermal coolingis provided by the surrounding tissue, which is not as efficient as thaton the surface. The result is that the ablation zone 278 has a largerdiameter D2 than D1, as determined by the heat transfer characteristicsof the surrounding tissue as well as the continued input of theultrasound energy from the beam 272. During this ultrasound-tissueinteraction, the ultrasound energy is being absorbed by the tissue 276,and less of it is available to travel further into the tissue. Thus acorrespondingly smaller diameter heated zone is developed in the tissue276, and the overall result is the formation of the heated ablation zone278 which is in the shape of an elongated tear drop limited to a depth288 into the tissue 276.

The formation of the ablation zone (including the size of the ablationzone and other characteristics) is dependent on time, as shown in FIGS.7A-7D, which show the formation of the lesion at times t1, t2, t3 andt4, respectively. As the sound beam 272 initially impinges on the frontsurface 280 of the tissue 276 at time t1, heat is created which beginsto form the lesion 278 (FIG. 7A). As time passes on to t2 and t3 (FIGS.7B and 7C), the ablation zone 278 continues to grow in diameter anddepth. This time sequence from t1 to t3 takes as little as about 1 to 5seconds, or preferably about 3 to 5 seconds, depending on the ultrasoundenergy density. As the incidence of the ultrasound beam 272 is continuedbeyond time t3, the ablation lesion 278 grows slightly in diameter andlength, and then stops growing due to the steady state achieved in theenergy transfer from its ultrasound form to the thermal form balanced bythe dissipation of the thermal energy into the surrounding tissue. Theexample shown in of FIG. 7D shows the lesion after an exposure t4 ofapproximately 30 seconds to the ultrasound beam 272. Thus the lesionreaches a natural limit in size and does not grow indefinitely.

The shape of the lesion or ablation zone 278 formed by the ultrasoundbeam 272 depends on factors such as the ultrasound beam 272, thetransducer element 3100 (including the material, the geometry, theportions of the transducer element 3100 that are energized and/or notenergized, etc.), any matching layers and/or backings present, theelectrical signal from the source of electrical energy (including thefrequency, the voltage, the duty cycle, the length and shape of thesignal, etc.), and the duration of energy delivery. The characteristicsof the target tissue include the thermal transfer properties and theultrasound absorption, attenuation, and backscatter properties of thetarget tissue and surrounding tissue. The size and characteristics ofthe ablation zone 278 also depend on the frequency and voltage appliedto the transducer element 3100 to create the desired ultrasound beam.

As mentioned above, properties such as the shape and construction of atransducer element influence the ablation lesions created by thetransducer element. The particular example lesion shown in FIGS. 7Athrough 7D is a tear-shaped lesion, for example as produced by atransducer element 3100 comprising a circular disc. A second variationof ablation shape is shown in FIG. 8, where the ablation zone 278′ has ashorter depth 288′. In this variation, the lesion 278′ has a more bluntshape than the ablation zone 278 of FIG. 6. One possible lesion geometryof this second variation may be a tooth-shaped geometry, as shown inFIG. 8, though the geometry may alternatively have a blunted tear shape,a circular shape, or an elliptical shape. As shown in FIG. 8, zone 278′(similarly to zone 278 in FIG. 6) has a diameter Dl smaller than thediameter D of the beam 272′ at the tissue surface 280 due to the thermalcooling provided by the surrounding fluid flowing past the tissuesurface 280. This variation in lesion geometry is produced by atransducer 3100 a having an acoustically inactive portion 4200 locatedat its center, i.e., a doughnut-shaped transducer which emits anultrasound beam 272′ that is generally more diffused, with a broader,flatter profile, than the ultrasound beam 272 shown in FIG. 6. Theultrasound beam 272′ emitted from such a doughnut-shaped transducer, asshown in FIG. 8, has reduced peak intensity along the midline of theenergy beam (as shown in cross section by the dotted lines in FIG. 8).With this ultrasound-tissue interaction, the reduced peak intensityalong the midline of the energy beam is absorbed by the tissue, and lessand less of the energy is available to travel further into the tissue,thereby resulting in a blunter lesion as compared to the firstvariation.

The ultrasound energy density determines the speed at which the ablationoccurs. The acoustic power delivered by the transducer element 3100,divided by the cross sectional area of the beamwidth, determines theenergy density per unit time. In the present embodiments, effectiveacoustic power ranges preferably from 0.5 to 25 watts, more preferablyfrom 2 to 10 watts, and even more preferably from 2 to 7 watts. Thecorresponding power densities range from approximately 50 watts/cm² to2500 watts/cm²). These power densities are developed in the ablationzone. As the beam diverges beyond the ablation zone, the energy densityfalls such that ablation will not occur, regardless of exposure time.

The transducer subassembly 3000 may additionally be coupled to a sensor(not shown). One variation of a sensor is a temperature sensor. Thetemperature sensor functions to detect the temperature of thesurrounding environment, the transducer element 3100, and/or thetemperature of any other suitable element or area. The sensor may alsobe used to monitor temperature of cooling fluid as it flows past thetransducer. The temperature sensor is a thermocouple, but mayalternatively be any suitable temperature sensor, such as a thermistoror an infrared temperature sensor. Optionally, the temperature sensor iscoupled to the transducer, for example, on the proximal face.Temperature information gathered by the sensor is used to manage thedelivery of continuous ablation energy to the tissue 276 during therapy,as well as to manage the temperature of the target tissue and/or theultrasound delivery system. In one embodiment, the sensor has a geometrythat is substantially identical to the geometry of the transducerelement 3100, so that the area diagnosed by the sensor is substantiallyidentical to the area to be treated by the transducer element 3100.Alternatively, the sensor has a smaller geometry to minimize interferingwith the delivery of ultrasound energy, but may be located in a regionthat is a local hot spot. For example, a small thermocouple mounted inthe center of the proximal heat spreader 3200 monitors the temperatureat the hottest spot of the transducer assembly. Additional details ontemperature sensors are disclosed in applications previouslyincorporated by reference above.

Alternatively, in a second variation of a sensor, the same ultrasoundtransducer element 3100 serves as a sensor and is used for the purposeof tissue detection. On the one hand, in order to achieve ablation, thetransducer element 3100 is used to generate and deliver an ultrasoundbeam of sufficient energy to the tissue in a manner such that the energyinput exceeds the thermal relaxation provided by the cooling due to thesurrounding tissue. This mode of energizing the ultrasound transducerelement 3100 is termed as the ablation mode. On the other hand, thetransducer element 3100 may be used to image tissue or to detect tissuecharacteristics, by utilizing an ultrasound signal optimized for tissuesensing which is generally not sufficient for heating of the tissue. Onesuch ultrasound imaging technique is referred to in the art as A-Mode,or Amplitude Mode imaging. This mode of energizing the transducerelement 3100 is termed as the imaging mode. The imaging mode is utilizedin directing the therapy provided by the ablation of the tissue. Thetransducer element 3100 can be used in the imaging mode in order todetect the gap (namely, the distance of the tissue surface from thedistal tip of the catheter 2000), the thickness of the tissue targetedfor ablation, characteristics of the ablated tissue, the incident beamangle, or any other suitable parameter or characteristic of the tissueand/or the environment around the ultrasound delivery system, such astemperature, thickness and ablation depth. Additional details on theseand other applicable features are described in the disclosurespreviously incorporated by reference.

Additionally and optionally, the ultrasound delivery system of thepreferred embodiments includes a processor, coupled to the sensor, thatcontrols the electrical attachments and/or the electrical signaldelivered to the electrical attachments, based on the informationobtained by the sensor. The processor may be a conventional processor,or it may alternatively be any suitable device to perform the desiredprocessing functions.

The processor receives information from the sensor, such as informationrelated to the distance between the catheter and the tissue (i.e., thegap distance), the thickness of the tissue targeted for ablation, thecharacteristics of the ablated tissue, or any other suitable parameteror characteristic. Based on this information, the processor controls theultrasound beam emitted by the transducer element 3100 by modifying theelectrical signal sent to the transducer element 3100 via the electricalattachment. This may include modifying the frequency, the voltage, theduty cycle, the length of the pulse, and/or any other suitableparameter. The processor may also control the ultrasound beam inmulti-element transducers by controlling which portions of thetransducer element are energized, and/or by controlling the frequency,voltage, duty cycle, etc. at which various portions of the transducerelement may be energized. Additionally, the processor may further becoupled to a fluid flow controller. The processor may control the fluidflow controller in order to increase or decrease fluid flow based on thedetected characteristics of the ablated tissue, of the unablated ortarget tissue, the temperature of the cooling fluid, tissue and/orenergy source, and/or any other suitable conditions. Further, theprocessor may control the fluid flow controller in order to maintain thetransducer element 3100 within a desired operating range oftemperatures. Further, the motion of the transducer to create a lesionline or shape in the tissue may be controlled either by an operator orvia one or more motors under processor control.

By controlling the ultrasound beam and/or the cooling of the targetedtissue or transducer element 3100, the shape of the ablation zone 278can be controlled. For example, the depth 288 of the ablation zone canbe controlled such that a transmural or substantially transmural lesionis achieved. Further, the nature of the lesion can be controlled bycontrolling the speed of the beam. The speed at which the beam movesalong the tissue determines the amount of energy deposited in thetissue. Thus, for example, slower speeds result in longer dwell times,thereby increasing the energy transferred to the tissue and, hence,creating deeper lesions. Additionally, the processor functions tominimize the possibility of creating a lesion beyond the targetedtissue, for example, beyond the outer atrial wall. If the sensor detectsthat the lesion and/or the ablation window is about to extend beyond theouter wall of the atrium, or that the depth of the lesion has reached orexceeded a preset depth, the processor turns off the power generatorand/or ceases to send electrical signals to the transducer and/or movesthe beam.

Additionally, the processor may function to maintain a preferred gapdistance between the transducer and the surface of the target tissue.The gap distance is preferably between 2 mm and 25 mm, more preferablybetween 2 mm and 20 mm, and even more preferably between 2 mm and 15 mm.If the sensor detects that the lesion and/or the ablation window isabout to extend beyond the outer wall of the atrium or is not reachingthe outer wall of the atrium, or that the depth of the lesion has notreached or has exceeded a preset depth, the processor may reposition theenergy delivery system. For example, as the catheter 2000 is rotated,the ablation window sweeps an ablation path (such as a circular orelliptical ablation path) creating a section of a conical shell.However, if the sensor determines that the ablation window is notreaching the wall of the atrium, the processor may move the elongatemember forwards or backwards along the Z-axis, or indicate that itshould be moved, in order to adjust for possible variations in anatomy.In such an embodiment, the operator can reposition the catheter 2000, orthe processor may be coupled to a motor drive unit or other control unitthat functions to position the catheter 2000.

While the above transducer elements and transducer subassemblies havebeen described in the context of ablation catheters, it should beunderstood that the transducer elements and transducer subassembliesdescribed herein can be used as part of any device configured toultrasonically image and/or ablate tissue. Additionally, while the aboveis a complete description of the preferred embodiments of the invention,various alternatives, modifications, and equivalents may be used.Therefore, the above description should not be taken as limiting thescope of the invention which is defined by the appended claims.

1.-25. (canceled)
 26. A method for creating a transmural lesion intissue, the method comprising: positioning a distal portion of acatheter near the tissue, wherein an ultrasound transducer is attachedto the distal portion and is operatively coupled to a console andprocessor; imaging the tissue by energizing the ultrasound transducer ata first power level to produce an ultrasound beam, wherein the imagingdetermines a thickness of the tissue, and a gap distance between theultrasound transducer and the tissue; ablating the tissue by energizingthe ultrasound transducer at a second power level to produce theultrasound beam; and controlling energy delivered to the tissue duringthe ablating, using the processor, wherein the processor adjusts a speedof the ultrasound beam moving across the tissue based on the thicknessand gap distance, to create the transmural lesion.
 27. The method ofclaim 26, wherein the ultrasound transducer is energized at a firstfrequency level during the imaging and a second frequency level whileablating.
 28. The method of claim 26, wherein there is no direct contactbetween the ultrasound transducer and the tissue during imaging or theablating.
 29. The method of claim 26, wherein the transmural lesion hasa tear drop shape.
 30. The method of claim 26, wherein during theablating, the transmural lesion is created while an outer layer oftissue remains substantially undamaged.
 31. The method of claim 26,wherein during the ablating, the transmural lesion is created while anendocardial surface of the tissue remains uncharred.
 32. The method ofclaim 26, wherein the ultrasound transducer comprises a single, flattransducer element.
 33. The method of claim 32, wherein the ultrasoundtransducer comprises an inactive portion surrounded by an activeportion.
 34. The method of claim 32, wherein a heat sink is attached tothe ultrasound transducer.
 35. The method of claim 32, wherein amatching layer is coupled to the ultrasound transducer.
 36. The methodof claim 26, wherein the processor further adjusts a catheter positionto maintain a preferred value for the gap distance during the ablating.37. The method of claim 26, wherein the processor further adjusts acatheter position to achieve a preset depth for the transmural lesion.38. The method of claim 26, wherein the processor indicates to anoperator to reposition the catheter for the ablating.
 39. The method ofclaim 26, further comprising supplying a cooling fluid in fluidcommunication with the ultrasound transducer.
 40. The method of claim39, wherein a temperature sensor is coupled to the ultrasound transducerfor monitoring temperature of the ultrasound transducer.
 41. The methodof claim 39, wherein the processor controls flow of the cooling fluid tomaintain a temperature of the ultrasound transducer within a desiredoperating range.
 42. The method of claim 39, wherein the processorcontrols flow of the cooling fluid based on detected characteristics ofthe tissue.
 43. The method of claim 26, wherein the processor, based onthe thickness and the gap distance, controls the ultrasound beam bymodifying an electrical signal sent to the ultrasound transducer. 44.The method of claim 26, wherein the processor uses the speed, the energydelivered, and a beamwidth of the ultrasound beam to deliver a desiredenergy density to the tissue.
 45. The method of claim 26, wherein theimaging further comprises determining an incident beam angle, and theincident beam angle is used to determine the speed of the ultrasoundbeam during ablation.